Ct detector having a segmented optical coupler and method of manufacturing same

ABSTRACT

The present invention is a directed to a CT detector for a CT imaging system that incorporates a segmented optical coupler between a photodiode array and a scintillator array. The segmented optical coupler also operates as a light collimator which improves the light collection efficiency of the photodiode array. The segmented optical coupler is defined by a series of reflector elements that collectively form a plurality of open cells. The open cells form light transmission cavities and facilitate the collimation of light from a scintillator to a photodiode. The cavities may be filled with optical epoxy for sealing to the photodiode array.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application is continuation of and claims priority of U.S.Ser. No. 11/163,973 filed Nov. 4, 2005, which is a divisional of andclaims priority of U.S. Ser. No. 10/908,209 filed May 2, 2005, now U.S.Pat. No. 7,064,334, which is a continuation of and claims priority ofU.S. Ser. No. 10/249,052 filed Mar. 12, 2003, now U.S. Pat. No.6,933,504.

BACKGROUND OF THE INVENTION

The present invention relates generally to diagnostic imaging and, moreparticularly, to a CT detector having a segmented or non-contiguousoptical coupler and method of manufacturing same. Additionally, thesegmented optical coupler operates as a light collimator integrallyformed between the scintillators and photodiodes of the detector.

Typically, in computed tomography (CT) imaging systems, an x-ray sourceemits a fan-shaped beam toward a subject or object, such as a patient ora piece of luggage. Hereinafter, the terms “subject” and “object” shallinclude anything capable of being imaged. The beam, after beingattenuated by the subject, impinges upon an array of radiationdetectors. The intensity of the attenuated beam radiation received atthe detector array is typically dependent upon the attenuation of thex-ray beam by the subject. Each detector element of the detector arrayproduces a separate electrical signal indicative of the attenuated beamreceived by each detector element. The electrical signals aretransmitted to a data processing system for analysis which ultimatelyproduces an image.

Generally, the x-ray source and the detector array are rotated about thegantry within an imaging plane and around the subject. X-ray sourcestypically include x-ray tubes, which emit the x-ray beam at a focalpoint. X-ray detectors typically include a collimator for collimatingx-ray beams received at the detector, a scintillator for convertingx-rays to light energy adjacent the collimator, and photodiodes forreceiving the light energy from the adjacent scintillator and producingelectrical signals therefrom.

Typically, each scintillator of a scintillator array converts x-rays tolight energy. Each scintillator discharges light energy to a photodiodeadjacent thereto. Each photodiode detects the light energy and generatesa corresponding electrical signal. The outputs of the photodiodes arethen transmitted to the data processing system for image reconstruction.

“Cross talk” between detector cells of a CT detector is common. “Crosstalk” is generally defined as the communication of data between adjacentcells of a CT detector. Generally, cross talk is sought to be reduced ascross talk leads to artifact presence in the final reconstructed CTimage and contributes to poor spatial resolution. Typically, fourdifferent types of cross talk may result within a single CT detector.X-ray cross talk may occur due to x-ray scattering between scintillatorcells. Optical cross talk may occur through the transmission of lightthrough the reflectors that surround the scintillators. Known CTdetectors utilize a contiguous optical coupling layer(s), typicallyepoxy, to secure the scintillator array to the photodiode array. Crosstalk, however, can occur as light from one cell is passed to anotherthrough the contiguous layer. Electrical cross talk can occur fromunwanted communication between photodiodes. Of the above types of crosstalk, cross talk though the contiguous optical coupler layer(s) isgenerally considered a major source of cross talk in the CT detector.

Therefore, it would be desirable to design a CT detector having improvedoptical coupling between the scintillator array and photodiode array toreduce cross talk in the CT detector and improve spatial resolution ofthe final reconstructed image.

BRIEF DESCRIPTION OF THE INVENTION

The present invention is a directed to a CT detector for a CT imagingsystem that overcomes the aforementioned drawbacks. The CT detectorincorporates a gridded light collimator between a photodiode array and ascintillator array. The light collimator improves the light collectionefficiency of the photodiode array and may be formed of reflectormaterial so as to reduce cross talk within the detector. Each griddedcollimator is defined by a series of reflector elements thatcollectively form a plurality of open cells. The open cells form lighttransmission cavities and facilitate the collimation of light from ascintillator to a photodiode. The cavities may be filled with opticalepoxy for sealing to the photodiode array or scintillator array therebyavoiding the drawbacks associated with contiguous optical couplerlayers.

Therefore, in accordance with the present invention, a CT detectorincludes a plurality of scintillators arranged in an array to receivex-rays and output light in response to the received x-rays. A pluralityof light detection elements are arranged in an array dimensionallysimilar to the scintillator array and are configured to detect lightfrom the scintillators. A non-contiguous optical coupler is then used tosecure the plurality of scintillators to the plurality of lightdetection elements.

According to another aspect of the present invention, a CT systemincludes a rotatable gantry having a bore centrally disposed therein anda table movable fore and aft through the bore and configured to positiona subject for CT data acquisition. A high frequency electromagneticenergy projection source is positioned within the rotatable gantry andconfigured to project high frequency electromagnetic energy toward thesubject. The CT system further includes a detector array disposed withinthe rotatable gantry and configured to detect high frequencyelectromagnetic energy projected by the projection source and impingedby the subject. The detector array includes a plurality of scintillatorsarranged in a scintillator array as well as a plurality of photodiodesarranged in a photodiode array. A light collimator having a plurality oflight transmission cavities is disposed between the scintillator arrayand the photodiode array.

In accordance with a further aspect of the present invention, a methodof CT detector manufacturing includes the steps of forming ascintillator array having a plurality of scintillators and forming aphotodiode array having a plurality of photodiodes. An open-celledcollimator is then deposited between the arrays. The resulting assemblyis then secured to one another.

Various other features, objects and advantages of the present inventionwill be made apparent from the following detailed description and thedrawings.

BRIEF DESCRIPTION OF DRAWINGS

The drawings illustrate one preferred embodiment presently contemplatedfor carrying out the invention.

In the drawings:

FIG. 1 is a pictorial view of a CT imaging system.

FIG. 2 is a block schematic diagram of the system illustrated in FIG. 1.

FIG. 3 is a perspective view of one embodiment of a CT system detectorarray.

FIG. 4 is a perspective view of one embodiment of a detector.

FIG. 5 is illustrative of various configurations of the detector in FIG.4 in a four-slice mode.

FIG. 6 is a schematic of a cross-section of a CT detector in accordancewith the present invention.

FIGS. 7-10 set forth steps of various techniques of manufacturing a CTdetector in accordance with the present invention.

FIG. 11 is a pictorial view of a CT system for use with a non-invasivepackage inspection system.

DETAILED DESCRIPTION

The operating environment of the present invention is described withrespect to a four-slice computed tomography (CT) system. However, itwill be appreciated by those skilled in the art that the presentinvention is equally applicable for use with single-slice or othermulti-slice configurations. Moreover, the present invention will bedescribed with respect to the detection and conversion of x-rays.However, one skilled in the art will further appreciate that the presentinvention is equally applicable for the detection and conversion ofother high frequency electromagnetic energy. The present invention willbe described with respect to a “third generation” CT scanner, but isequally applicable with other CT systems.

Referring to FIGS. 1 and 2, a computed tomography (CT) imaging system 10is shown as including a gantry 12 representative of a “third generation”CT scanner. Gantry 12 has an x-ray source 14 that projects a beam ofx-rays 16 toward a detector array 18 on the opposite side of the gantry12. Detector array 18 is formed by a plurality of detectors 20 whichtogether sense the projected x-rays that pass through a medical patient22. Each detector 20 produces an electrical signal that represents theintensity of an impinging x-ray beam and hence the attenuated beam as itpasses through the patient 22. During a scan to acquire x-ray projectiondata, gantry 12 and the components mounted thereon rotate about a centerof rotation 24.

Rotation of gantry 12 and the operation of x-ray source 14 are governedby a control mechanism 26 of CT system 10. Control mechanism 26 includesan x-ray controller 28 that provides power and timing signals to anx-ray source 14 and a gantry motor controller 30 that controls therotational speed and position of gantry 12. A data acquisition system(DAS) 32 in control mechanism 26 samples analog data from detectors 20and converts the data to digital signals for subsequent processing. Animage reconstructor 34 receives sampled and digitized x-ray data fromDAS 32 and performs high speed reconstruction. The reconstructed imageis applied as an input to a computer 36 which stores the image in a massstorage device 38.

Computer 36 also receives commands and scanning parameters from anoperator via console 40 that has a keyboard. An associated cathode raytube display 42 allows the operator to observe the reconstructed imageand other data from computer 36. The operator supplied commands andparameters are used by computer 36 to provide control signals andinformation to DAS 32, x-ray controller 28 and gantry motor controller30. In addition, computer 36 operates a table motor controller 44 whichcontrols a motorized table 46 to position patient 22 and gantry 12.Particularly, table 46 moves portions of patient 22 through a gantryopening 48.

As shown in FIGS. 3 and 4, detector array 18 includes a plurality ofscintillators 57 forming a scintillator array 56. A collimator (notshown) is positioned above scintillator array 56 to collimate x-raybeams 16 before such beams impinge upon scintillator array 56.

In one embodiment, shown in FIG. 3, detector array 18 includes 57detectors 20, each detector 20 having an array size of 16×16. As aresult, array 18 has 16 rows and 912 columns (16×57 detectors) whichallows 16 simultaneous slices of data to be collected with each rotationof gantry 12.

Switch arrays 80 and 82, FIG. 4, are multi-dimensional semiconductorarrays coupled between scintillator array 56 and DAS 32. Switch arrays80 and 82 include a plurality of field effect transistors (FET) (notshown) arranged as multi-dimensional array. The FET array includes anumber of electrical leads connected to each of the respectivephotodiodes 60 and a number of output leads electrically connected toDAS 32 via a flexible electrical interface 84. Particularly, aboutone-half of photodiode outputs are electrically connected to switch 80with the other one-half of photodiode outputs electrically connected toswitch 82. Additionally, a reflector layer (not shown) may be interposedbetween each scintillator 57 to reduce light scattering from adjacentscintillators. Each detector 20 is secured to a detector frame 77, FIG.3, by mounting brackets 79.

Switch arrays 80 and 82 further include a decoder (not shown) thatenables, disables, or combines photodiode outputs in accordance with adesired number of slices and slice resolutions for each slice. Decoder,in one embodiment, is a decoder chip or a FET controller as known in theart. Decoder includes a plurality of output and control lines coupled toswitch arrays 80 and 82 and DAS 32. In one embodiment defined as a 16slice mode, decoder enables switch arrays 80 and 82 so that all rows ofthe photodiode array 52 are activated, resulting in 16 simultaneousslices of data for processing by DAS 32. Of course, many other slicecombinations are possible. For example, decoder may also select fromother slice modes, including one, two, and four-slice modes.

As shown in FIG. 5, by transmitting the appropriate decoderinstructions, switch arrays 80 and 82 can be configured in thefour-slice mode so that the data is collected from four slices of one ormore rows of photodiode array 52. Depending upon the specificconfiguration of switch arrays 80 and 82, various combinations ofphotodiodes 60 can be enabled, disabled, or combined so that the slicethickness may consist of one, two, three, or four rows of scintillatorarray elements 57. Additional examples include, a single slice modeincluding one slice with slices ranging from 1.25 mm thick to 20 mmthick, and a two slice mode including two slices with slices rangingfrom 1.25 mm thick to 10 mm thick. Additional modes beyond thosedescribed are contemplated.

Referring now to FIG. 6, a schematic of a cross-section of a CT detector20 is shown. As discussed above, detector 20 includes a scintillatorarray 56 defined by a plurality of scintillators 57. Each of thescintillators is designed to generate a light output 85 in response thereception of x-rays 16. A reflector layer 86 coats the x-ray receptionsurface of the scintillators to improve light collection efficiency ofthe photodiodes. The reflector layer 86 is composed of a material thatallows x-rays projected from a projection source to pass through andreflects light generated by the scintillators back toward thephotodiodes. The reflector layer is integrated with a series ofreflector elements 88 that extend between adjacent scintillators 57 as areflector wall. The reflector elements 88 are designed to prevent lightscattering and/or reduce x-ray scattering between scintillators.

CT detector 20 is constructed such that a light cavity 90 extendsbetween each photodiode and scintillator. The light cavity may beconstructed in accordance with a number of fabrication techniques aswill be described with respect to FIGS. 7-10 and is defined by cavityelements or plates 92. Plates 92 are preferably formed of a reflectormaterial similar to that used to form reflector elements 88.Additionally, plates 92 have a width similar to the width of thereflector elements 88. Preferably, plates 92 are formed during theformation of reflector elements 88, as will be described with respect toFIG. 7. As such, plates 92 extend from reflector elements to the lightdetection surface of the photodiode array.

Plates 92 are constructed to form light transmission cavities 90 and, assuch, operate as an inner-cell light collimator. Plates 92 are designedto eliminate light cross talk between scintillators thereby collimatinglight toward the light detection surfaces of the photodiode array.Further, plates 92 may be coated with an optical coupling film or resinso as to secure the plates to the photodiode array. Alternately, theplates may be bonded to the surface of the photodiode array. In afurther embodiment, each of the light transmission cavities 90 is filledwith an optical epoxy similar to the epoxy used in a contiguous epoxylayer. The optical epoxy operates as adhesive to connect the photodiodearray to the scintillator array. With the presence of reflector plates92, the drawbacks associated with contiguous optical layer cross talkare avoided. While epoxy may be used to secure the arrays to oneanother, other composites and materials such as thermoplastics may beused and are within the scope of the invention.

Referring now to FIG. 7, steps for a technique of manufacturing a CTdetector similar to that described with respect to FIG. 6 are shown. Thesteps illustrated may be carried out by a labor intensive process, afully automated, computer driven process, or a combination thereof.Technique 100 begins at 102 with the assimilation of products, personneland the like for CT detector fabrication. That achieved during this stepmay vary but, at a minimum, should include the preparation of ascintillator block. The scintillator block is then mounted onto adissolvable material 104. The scintillator block and dissolvablematerial are then diced or cut at 106. Once cut, either along one or twodimensions, a plurality of scintillator cells uniformly spaced from oneanother results. Reflector material is then cast at 108 in the voidscreated between the scintillator cells as a result of the dicingprocess. The reflector material should be cast such that the interfacebetween scintillators is completely filled as is the interface betweenadjacent portions of the dissolvable material. The cast reflectormaterial is then allowed to cure and undergoes any additional processingto insure proper reflectivity and the like. Once the cast reflectormaterial has cured, the dissolvable material is dissolved at 110. Theprocess for dissolving the material depends on the type of dissolvablematerial used. For example, the dissolvable material may be placed in awash and chemically dissolved or heated at a specified temperature to,in essence, “melt” away the dissolvable material. After the dissolvingprocess is complete, a scintillator array with an integrated castreflector results. Of particular note is that each reflector elementbetween the scintillators extends beyond the scintillator, i.e. has agreater length than the scintillators. The portion of the reflector thatextends beyond the scintillator operates as a reflector plate asdescribed above. The open cells that result between reflector platesdefine a light transmission cavity and are filled with optical epoxy at112. The optical epoxy permits the transmission of light betweenscintillator and photodiode while simultaneously creating an adhesioninterface for coupling the scintillator to the photodiode. As such, thephotodiode array and scintillator array are coupled to one another at114. This portion of the CT detector fabrication process is thencomplete and the remainder of the CT detector fabricating takes placedownstream at 116.

The CT detector described with respect to FIG. 6 and fabricatedaccording to the technique of FIG. 7 illustrates only one example of thepresent invention. A similar CT detector incorporating the advantages ofthat described with respect to FIG. 6 and fabricated in accordance withtechniques different from that illustrated in FIG. 7 are contemplatedand within the scope of this invention. For purpose of illustration andnot limitation, additional manufacturing techniques and the resultingstructures will be described with reference to FIGS. 8-10.

Referring now to FIG. 8, another CT manufacturing process 118 begins at120 with a block of scintillator material being prepared. The block isthen placed onto a block of thermoplastic material at 122. Thescintillator block and the thermoplastic are then diced or cut 124 inaccordance with known dicing processes. Preferably, only a portion ofthe thermoplastic is diced thereby leaving a thin, uncut portion thatcan be used to seal against the photodiode array. Cast reflector is thendeposited in the voids 126 between scintillator cells that result fromthe dicing process. In contrast to the CT detector constructed inaccordance with FIG. 7, an optical epoxy between the reflector platesformed by the cast reflector is not used. Because the thermoplasticmaterial is not completely diced through, a thin thermoplastic layerresults that, as discussed above, is used to secure the scintillatorarray to the photodiode array as opposed to an optical epoxy. Process118 then concludes at 128 with the CT detector undergoing additionalprocessing and fabrication in accordance with known techniques.

The processes described above involve alternations to the scintillatorarray. In contrast, the process of FIG. 9 creates the reflector platesby etching the photodiode array. Specifically, process 130 begins at 132with the formation of a photodiode array. At 134, the photodiode arrayis coated with a film of semiconductor or other suitable materials.Preferably, a thin layer of Silicon is applied or thermally grown andallowed to cure to the photodiode light reception surface. Semiconductormaterials that will not adversely affect the light collection abilitiesof the photodiode array should be used. The surface of the photodiodearray is then masked and plasma etched at 136 using standardsemiconductor fabrication techniques to form a grid. Varioussemiconductor fabrication processes are contemplated including chemicaletching, mechanical etching, ion beam milling, and the like. The resultof the etching process should result in a series of open cells definedby the semiconductor material. The open cells should be verticallyaligned with the light detection surfaces of the photodiode array. Theopen cells are then filled with optical epoxy 138 to secure thephotodiode array to the scintillator array at 140. The resultingassembly then undergoes standard post-processing techniques whereuponthe process ends at 142.

The process illustrated in FIG. 10 utilizes an intermediary element thatis not integrated with the scintillator array or photodiode array.Manufacturing process 144 begins at 146 with the formation of ascintillator array and a photodiode array in accordance with knownfabrication techniques. A grid is then etched at 148 from a sheet ofthin metallic or other material. The grid defines a number of cellsdimensionally equivalent to the scintillators and photodiodes.Additionally, the grid preferably has a height equal to the desiredheight of the light transmission cavities heretofore described.Accordingly, the open cells formed in the grid are aligned with theeither the scintillators of the scintillator array or with thephotodiodes of the photodiode array at 150. The grid is then bonded at152 to the selected array. The open cells or cavities defined by thegrid may then be filled with optical epoxy at 154. The optical epoxy isthen used to secure the selected array to the other array at 156.Alternately, the open cells may be left empty and the grid bonded to theother array. The process is then complete at 158.

Each of the above-described manufacturing processes results in a CTdetector having a non-contiguous optical coupler thereby avoiding thedrawbacks associated with a contiguous optical coupler layer. Each ofthe processes produces a CT detector wherein a light transmission cavityis formed to collimate light emissions from a scintillator to aphotodiode. The cavity may be filled with optical coupling epoxy or leftempty and the scintillator bonded to the photodiode array. It ispreferred that the cavities be filled with epoxy as this results inbetter optical transmission and a stronger connection being formedbetween the scintillator and photodiode.

Referring now to FIG. 11, package/baggage inspection system 160 includesa rotatable gantry 162 having an opening 164 therein through whichpackages or pieces of baggage may pass. The rotatable gantry 162 housesa high frequency electromagnetic energy source 166 as well as a detectorassembly 168. A conveyor system 170 is also provided and includes aconveyor belt 172 supported by structure 174 to automatically andcontinuously pass packages or baggage pieces 176 through opening 164 tobe scanned. Objects 176 are fed through opening 164 by conveyor belt172, imaging data is then acquired, and the conveyor belt 172 removesthe packages 176 from opening 164 in a controlled and continuous manner.As a result, postal inspectors, baggage handlers, and other securitypersonnel may non-invasively inspect the contents of packages 176 forexplosives, knives, guns, contraband, etc.

Therefore, in accordance with one embodiment of the present invention, aCT detector includes a plurality of scintillators arranged in an arrayto receive x-rays and output light in response to the received x-rays. Aplurality of light detection elements are arranged in an arraydimensionally similar to the scintillator array and are configured todetect light from the scintillators. A non-contiguous optical coupler isthen used to secure the plurality of scintillators to the plurality oflight detection elements.

According to another embodiment of the present invention, a CT systemincludes a rotatable gantry having a bore centrally disposed therein anda table movable fore and aft through the bore and configured to positiona subject for CT data acquisition. A high frequency electromagneticenergy projection source is positioned within the rotatable gantry andconfigured to project high frequency electromagnetic energy toward thesubject. The CT system further includes a detector array disposed withinthe rotatable gantry and configured to detect high frequencyelectromagnetic energy projected by the projection source and impingedby the subject. The detector array includes a plurality of scintillatorsarranged in a scintillator array as well as a plurality of photodiodesarranged in a photodiode array. A light collimator having a plurality oflight transmission cavities is disposed between the scintillator arrayand the photodiode array.

In accordance with a further embodiment of the present invention, amethod of CT detector manufacturing includes the steps of forming ascintillator array having a plurality of scintillators and forming aphotodiode array having a plurality of photodiodes. An open-celledcollimator is then deposited between the arrays. The resulting assemblyis then secured to one another.

The present invention has been described in terms of the preferredembodiment, and it is recognized that equivalents, alternatives, andmodifications, aside from those expressly stated, are possible andwithin the scope of the appending claims.

1. A CT system, comprising: a rotatable gantry having a bore centrally disposed therein; a table movable fore and aft through the bore and configured to position a subject for CT data acquisition; a high frequency electromagnetic energy projection source positioned within the rotatable gantry and configured to project high frequency electromagnetic energy toward the subject; and a detector array disposed within the rotatable gantry and configured to detect high frequency electromagnetic energy projected by the projection source and impinged by the subject, wherein the detector array comprises: a plurality of scintillators arranged in a scintillator array; a plurality of photodiodes arranged in a photodiode array; and a 2D gridded light collimator plate array having a plurality of collimator plates that collectively define a plurality of light transmission cavities disposed between the plurality of scintillators and the plurality of photodiodes, the plurality of collimator plates formed of light reflective material and arranged such that each light transmission cavity is disposed entirely between a light reception surface of a photodiode and a light emission surface of a scintillator.
 2. The CT system of claim 1, wherein the detector array includes a thin film optical coupler that secures the 2D gridded light collimator plate array to the photodiode array.
 3. The CT system of claim 2, wherein the thin film optical coupler comprises an optical coupling film or resin that coats the plurality of collimator plates and secures the plurality of collimator plates to the plurality of photodiodes.
 4. The CT system of claim 1, wherein the plurality of collimator plates is bonded to a light reception surface of the photodiode array.
 5. The CT system of claim 1, wherein the plurality of light transmission cavities is filled with an optical epoxy that permits transmission of light between the scintillator array and the photodiode array and operates as adhesive to secure the scintillator array to the photodiode array.
 6. The CT system of claim 1, wherein the plurality of light transmission cavities is filled with a thermoplastic that operates as adhesive to secure the photodiode array to the scintillator array.
 7. The CT system of claim 1, wherein the detector array comprises a plurality of reflector elements that extend as a reflector wall between adjacent instances of the plurality of scintillators, wherein the plurality of collimator plates and the plurality of reflector elements comprise a substantially same material.
 8. The CT system of claim 7, wherein the scintillator array is integrated with the plurality of reflector elements, wherein the plurality of collimator plates and the plurality of reflector elements are unitary.
 9. The CT system of claim 7, wherein the detector array comprises a reflector layer that coats a high frequency electromagnetic energy reception surface of the plurality of scintillators, wherein the plurality of reflector elements is integrated with the reflector layer; wherein the plurality of collimator plates, the plurality of reflector elements, and the reflector layer comprise the substantially same material, wherein the substantially same material allows the high frequency electromagnetic energy projected from the high frequency electromagnetic energy projection source to pass therethrough and reflects light generated by the plurality of scintillators toward the plurality of photodiodes.
 10. The CT system of claim 1, wherein the plurality of light transmission cavities comprises a series of cells defined by semiconductor material and vertically aligned with light emission surfaces of the scintillator array and/or light reception surfaces of the photodiode array.
 11. The CT system of claim 10, wherein the series of cells defined by the semiconductor material are filled with an optical epoxy that permits transmission of light between the scintillator array and the photodiode array and operates as adhesive to secure the scintillator array to the photodiode array.
 12. The CT system of claim 1, wherein the 2D gridded light collimator plate array comprises a grid that defines a plurality of cells dimensionally equivalent to the plurality of scintillators and the plurality of photodiodes.
 13. The CT system of claim 1 wherein each light transmission cavity collimates light between a single scintillator and a single photodiode.
 14. A CT scanner having a detector array designed to convert received x-rays to electrical signals that can be processed to reconstruct a CT image, the detector array comprising: a scintillator array; a photodiode array; and a light collimator array formed between the scintillator array and the photodiode array, the light collimator array arranged to sit atop a light reception surface of the photodiode array and define a single optical transmission window between an aligned scintillator and an aligned photodiode, wherein the light collimator array is defined by a 2D grid of collimator plates, each collimator plate including light reflective material.
 15. The CT scanner of claim 14, wherein the light collimator array is bonded to at least one of the scintillator array and the photodiode array.
 16. The CT scanner of claim 14, wherein each collimator plate further includes x-ray absorbent material.
 17. The CT scanner of claim 14, wherein voids defined between adjacent collimator plates are epoxy-filled.
 18. A CT detector, comprising: a plurality of scintillators arranged in an array to receive x-rays and output light in response to the reception of x-rays; a plurality of reflector elements with a single reflector element disposed between adjacent scintillators; a plurality of light detection elements arranged in an array to output electrical signals in response to light detected from the plurality of scintillators; and a 2D gridded light collimator plate optical coupler that permits transmission of light between the plurality of scintillators and the plurality of light detection elements and secures the plurality of scintillators to the plurality of light detection elements, wherein the 2D gridded light collimator plate optical coupler comprises a plurality of light transmission cavities defined by light collimator plates disposed between the plurality of scintillators and the plurality of light detection elements.
 19. The CT detector of claim 18, wherein the 2D gridded light collimator plate optical coupler comprises a non-contiguous optical coupler that defines the plurality of light transmission cavities between the plurality of scintillators and the plurality of light detection elements.
 20. The CT detector of claim 18, wherein the light collimator plates are arranged in an array that sits atop a light reception surface of the plurality of light detection elements and defines a single optical transmission window between an aligned scintillator and an aligned light detection element. 